Sensor and method for detection of a target substance

ABSTRACT

A sensor, in one configuration, includes a plurality of beads located in a channel. The beads are attached to a surface of the channel by a biological binding mechanism that can be selectively unbound in the presence of a target substance. A measurement system is configured and arranged to determine the population of beads in the channel after the introduction of an analyte solution into the channel.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 11/080,064, filed Mar. 14, 2005, entitled “A Sensor And MethodFor Detection Of A Target Substance”, which is a continuation-in-part ofU.S. patent application Ser. No. 10/944,140, filed Sep. 16, 2004,entitled “A Sensor And Method For Detection Of A Target Substance”,which claims priority under 35 U.S.C. 119(e) from Provisional U.S.patent application Ser. No. 60/504,334, filed Sep. 17, 2003, entitled“Electrochemical, Reagentless, Hand-Held Sensor And Method For DetectionOf DNA Hybridization And Other Molecular Binding And Cleaving Events”,both of which are incorporated herein by reference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

The techniques and mechanisms of the present invention were made withGovernment support under National Science Foundation Grant No.CHE0079225.

BACKGROUND

The modern concept of biological sensors evolved considerably over theforty years since it was first demonstrated that enzymes could beimmobilized at a surface of an electrochemical detector. Demands toreduce the sample volume, the cost and the time of analysis, and toincrease sensitivity and selectivity have been most pressing inbiological and biomedical sciences. This has fueled development ofmicro-analytical devices and bio-sensors with a wide range ofapplications in the clinical and defense settings, in gene and forensicanalysis, in environmental monitoring, food safety and many othersettings. While selectivity of biological sensors always derives fromthe unique molecular recognition interactions, such as antibody-antigenbinding or complimentary deoxyribonucleic acid (DNA) hybridization,transduction and amplification of these events into analytically usefulsignal is often a major challenge.

Interest in detecting DNA in a sequence-specific manner has grownsteadily in recent years. The ability to rapidly and inexpensivelydetect DNA of specific sequences may allow more efficient pathogen,point mutation and gene detection. Multiple approaches of detecting DNAin a sequence-specific fashion have been explored, including optical(for example, chemiluminescence fluorescence, Raman spectroscopic andsurface plasmon resonance), electronic as well as numerouselectrochemical methods. Historically, fluorescence methods have beenthe most sensitive. However, advances in the electrochemical detectionof DNA are becoming competitive in terms of sensitivity.

Detection of pathogenic species in water is also an important andchallenging problem. Several examples reporting direct response sensordevices sensitive to bacteria, viruses and bacterial toxins exist. Theseinclude, for example, colorimetric sensors designed to detect influenzavirus and E. Coli and an electrochemical sensor responding to E. Colienterotoxin.

These technologies, however, have certain disadvantages. It is thereforedesirable to provide improved methods and apparatus for electrochemicalsensing of a target substance. In one example, it is desirable toprovide sensors having microfabrication efficiencies that are highlyaccurate and yet inexpensive enough so that they are disposable.

SUMMARY

In one aspect, the invention features a sensor. The sensor comprisesmolecular tethers attached to a sensor surface such that at least someof a plurality of beads, as a result of a biological interaction betweena target substance and the tethers, can at least one of be attached tothe sensor surface by the tethers or released from the sensor surface bycleaving the tethers. The sensor further includes a measurement systemconfigured and arranged to determine the population of beads at thesensor surface after the introduction of the analyte into the sensor.

Various implementations of the invention may include one or more of thefollowing features. A tether includes a hybrid double-stranded DNAmolecule, an antibody, a modified antibody, a modified antigen or apeptide.

In yet another aspect, the invention is directed to a sensor todetermine if a target substance is present in an analyte. The sensorcomprises a plurality of beads located in a channel and attached to asurface of the channel by a molecular tether. The tether can beselectively cleaved in the presence of a target substance to release atleast some of the beads from the surface of the channel. The sensorfurther includes a measurement system configured and arranged todetermine the population of beads in the channel after the introductionof the analyte into the channel.

Various implementations of the invention may include one or more of thefollowing features. The target substance is a target antigen and amolecular tether includes a modified antigen and an antibody with themodified antigen having a lower binding constant than the targetantigen. The modified antigen has an approximately 2 orders of magnitudelower binding constant than the target antigen. The target substance isa target DNA and a molecular tether is a hybridized double-stranded DNAtether with one of the single-stranded DNA fragments of thedouble-stranded DNA tether being complementary to the target DNA. Thehybridized double-stranded DNA tether includes first and secondsingle-strand DNA fragments. The first single-strand DNA fragment isattached to a surface of the channel, and the sequence of the firstsingle-strand DNA fragment is complementary to the target DNA. Thesecond single-strand DNA fragment is attached to a surface of a bead.The measurement system may be a signal transduction system, aconductivity measurement system, a coulometric measurement system, aspectrometric measurement system, an optical measurement system, or anelectrochemical counting system.

In another aspect, the invention is directed to a sensor to detect atarget antigen. The sensor comprises a plurality of beads attached to asensor surface by a modified antigen and an antibody wherein themodified antigen has a lower binding constant than the target antigen.

In still another aspect, the invention is directed to a sensor to detecta target antibody. The sensor comprises a plurality of beads attached toa sensor surface by a modified antibody and an antigen wherein themodified antibody has a lower binding constant than the target antibody.

In another aspect, the invention is directed to a sensor to detect atarget substance in an analyte. The sensor comprises a channel and aplurality of beads are located in the channel and attached to a surfaceof the channel by a molecular tether. The tether can be selectivelycleaved in the presence of a target substance to release at least someof the beads from the surface of the channel. Measurement means areprovided for determining the population of beads in the channel afterthe introduction of the analyte into the channel.

In yet another aspect, the invention features a method of detecting atarget substance. The method includes providing a number of moleculartethers at a channel surface of a sensor such that at least some of aplurality of beads, as a result of a biological interaction between atarget substance and the tethers, can at least one of be attached to thechannel surface by the tethers or released from the channel surface bycleaving the tethers. An analyte is introduced into the channel.Thereafter, the population of beads at the channel surface is determinedto indicate a presence of the target substance in the analyte.

Various implementations of the invention may include one or more of thefollowing features. The tethers include a hybrid double-stranded DNAmolecule, an antibody, a modified antibody, a modified antigen or apeptide. The target substance is a target DNA and a molecular tether isa hybridized double-stranded DNA tether with one of the single-strandedDNA fragments of the double-stranded tether being complementary to thetarget DNA. The hybridized double-stranded DNA tether includes first andsecond single-strand DNA fragments. The first single-strand DNA fragmentis attached to a surface of the channel, and the sequence of the firstsingle-strand DNA fragment is complementary to the target DNA. Thesecond single-strand DNA fragment is attached to a surface of a bead.

Other implementations of the invention may include one or more of thefollowing features. The target substance is a target antigen and amolecular tether includes a modified antigen and an antibody with themodified antigen having a lower binding constant than the targetantigen. The target substance is a target antibody and a moleculartether includes a modified antibody and an antigen with the modifiedantibody having a lower binding constant than the target antibody.

The invention can include one or more of the following advantages. Thesensor can detect a target substance such as DNA and a range ofimmunogens (antigens), a large group of macromolecules includingcancer-specific proteins, toxins and numerous other pathogens. Thesensor may be microfabricated and made small enough to be hand-held. Thesensor may be disposable. The sensor functions without the use ofreagents. The sensor may be part of an array. Each sensor of the arraymay respond to a specific DNA sequence. The sensor can simultaneouslydetect different target substances in sub-micron liter (uL) volumes of asolution at nano-Molar (nM) concentration levels.

These and other features and advantages of the present invention will bepresented in more detail in the following specification of the inventionand the accompanying figures, which illustrate, by way of example, theprinciples of the invention.

BRIEF DESCRIPTION OF DRAWINGS

The invention may best be understood by reference to the followingdescription taken in conjunction with the accompanying drawings thatillustrate specific embodiments of the present invention.

FIG. 1A is a diagrammatic side view representation of a sensor accordingto the present invention.

FIG. 1B is a diagrammatic enlarged view of a portion of the sensor ofFIG. 1A.

FIG. 1C is a diagrammatic representation of the operation of the sensorof FIG. 1A.

FIG. 2A is a diagrammatic top view of a sensor device.

FIG. 2B is a diagrammatic enlarged view of a portion of the sensordevice of FIG. 2B.

FIG. 2C is a diagrammatic view of the sensor of FIG. 2A without anybeads in the sensor channel.

FIG. 3 graphically illustrates the reproductibility of potential versustime (E vs t) transients due to spectator species diffusion in a sensorchannel.

FIG. 4 is a diagrammatic representation of the operation of a sensor inaccordance with another embodiment of the present invention.

FIG. 5 is a diagrammatic representation of the operation of a sensor inaccordance with yet another embodiment of the present invention.

FIG. 6 is a diagrammatic representation of a system to monitormicro-bead detachment.

DETAILED DESCRIPTION

Reference will now be made in detail to some specific embodiments of theinvention including the best modes contemplated by the inventor forcarrying out the invention. Examples of these specific embodiments areillustrated in the accompanying drawings. While the invention isdescribed in conjunction with these specific embodiments, it will beunderstood that it is not intended to limit the invention to thedescribed embodiments. On the contrary, it is intended to coveralternatives, modifications, and equivalents as may be included withinthe spirit and scope of the invention as defined by the appended claims.For example, the techniques of the present invention will be describedin the context of a glass-based electrochemical sensor, although othermaterials, such as plastics and polymers, could also be used.

In the following description, numerous specific details are set forth inorder to provide a thorough understanding of the present invention. Thepresent invention may be practiced without some or all of these specificdetails. In other instances, well known process operations have not beendescribed in detail in order not to unnecessarily obscure the presentinvention.

As shown in FIGS. 1A and 1B, a sensor 10 includes an upper plate ormembrane 12 spaced from a lower plate or membrane 14 to form a channel16. A first electrode 18 is located at one end of the channel, and asecond electrode 20 is located at an opposite end of the channel. Thedistance between the electrodes is (g). The electrodes have a length(l). (See FIG. 2A).

The electrode 18 is a generator electrode. It generates a spectatorspecies, for example, chloride ions, that diffuse in an aqueous solutionin the channel 16, and which do not chemically interact with the anychemical species in the channel. The spectator species are generated ata constant rate galvanostically. In other words, to generate spectatorspecies such as chloride, a constant current pulse is applied to thegenerator electrode.

The second electrode 20 is a sensor electrode that potentiometricallymonitors the arrival rate of the spectator species. It is thusconfigured to detect the concentration of the spectator species at itsend of the channel, after the spectator species has traveled along thechannel from the generator electrode to the sensor electrode.

Both electrodes may be formed on, as shown, an inner surface 14 a of thelower plate 14. Alternatively, the electrodes could be fabricated on aninner surface 12 a of the upper plate 12. As discussed below, the sensorwould also include appropriate counter and reference electrodes.

The channel is populated by a plurality of beads or micro-beads 22. Thebeads also act as spacers in the channel between the top plate 12 andthe bottom plate 14. Alternatively, spacer dots, as discussed below, maybe used to space the upper plate 12 from the lower surface 14. The beadsmay be spherical in shape. Other shapes, however, are also possible. Forinstance, the beads may be cubical in shape.

The beads are attached to a surface of the channel by a molecular tether(molecular bridge) involving the target substance (for example, ssDNA orantigen). The molecular tether can be cleaved (broken) in the presenceof a target substance that the sensor is designed to detect. As shown inFIG. 1B, for example, the beads may be attached to the inner surface 14a of the bottom surface 14 by a number of DNA (deoxyribonucleic acid)tethers 24.

The tethers 24, in one embodiment, are formed by the hybridization oftwo single-strand DNA (ssDNA) fragments. As is known, DNA hybridizationoccurs when a single-stranded DNA of a particular sequence of nucleotidebases interacts with a “Watson-Crick” complementary strand of a ssDNA,resulting in the formation of a hybrid, double stranded DNA (dsDNA)molecule. One fragment 24 a, in the presently-described embodiment, isinitially attached to the surface 14 a of the membrane 14 in the gapbetween the generator electrode 18 and the sensor electrode 20. Thesequence of this ssDNA fragment is selected to be complementary to thetarget ssDNA, that is, the DNA to be detected by the sensor. The second,shorter ssDNA fragment 24 b is attached to the surface of the bead 22.The number of hybridized tethers linking each bead to the channelsurface 14 a is controlled by the surface density of the two components.The high bead density shown in FIG. 1A corresponds to the desiredinitial state of the sensor.

The functioning of the sensor involves two independent elements:recognition of a specific sequence of the target DNA, for example, and asignal transduction function. The signal transduction function involvesmeasurements of the rate of diffusion of the spectator species along thechannel in the inter-electrode gap, the gap between the electrodes 18and 20. The length of this gap, as noted, is (g). The diffusion of thenon-interacting spectator species is monitored by recording potentialversus time (E vs t) transients at the potentiometric sensor electrode20. The rise time and the shape of the transients depend on the freevolume of the solution in the channel between the generator electrode 18and the sensor electrode 20, and thus on the number and density of themicro-beads in the channel.

The E vs t transient recorded in the initial state of the sensor isexpected to exhibit a rapid rise of potential at the sensor electrode 20due to a rapid diffusion of the spectator species through the relativelysmall free volume in the channel. The latter can be computed bysubtracting the volume of the beads from the total volume of solution inthe channel. The theoretical expectation of the functional dependent ofthe measured sensor's potential on time is available by solving Fick'sequation of diffusion. The concentration of the spectator species versustime at the sensor electrode, a distance (g) from the generatorelectrode, C(g,t), obeys the following equation obtained by solvingFick's law of diffusion: $\begin{matrix}{{C\left( {g,t} \right)} = {C_{in} + {\frac{i}{{nF}\quad\rho\quad{AD}}\begin{Bmatrix}{{2\left( \frac{Dt}{\pi} \right)^{1/2}{\exp\left( {- \frac{g^{2}}{4{Dt}}} \right)}} -} \\{g \cdot {{erfc}\left\lbrack \frac{g}{2({Dt})^{1/2}} \right\rbrack}}\end{Bmatrix}}}} & (1)\end{matrix}$where C_(in) is the initial, background spectator species; A is thecross-sectional area of the channel (a product of the channel height (h)and the length of the electrodes (l)); ρ is a fractional free volume ofthe channel; D is the diffusion constant of spectator species; i is themagnitude of the current used to generate the spectator species; n isthe number of electrons involved in the elementary generation reaction;and F is the Faraday constant (approximately 96,500 C/mol).

The inter-electrode gap (g) may be between about 10 and 50 microns (um).Spherical beads may have a diameter of approximately 2-5 microns and adensity of about 2.5 g/cm³. The thickness or height of the channel (h)is determined by the diameter of the beads, and/or spacers of the sameheight. Thus, for a 2 micron bead, (h) would be equal to approximately 2microns. The length of the electrodes (1) may be between about 50 and1,000 microns. The diffusion of the spectator species between thegenerator and sensor electrodes is linear as the channel height is muchless than its length (h<<g).

The plates 12 and 14 may be made of an electrically insulating glass,such as microscope slides. Alternatively, the plates may be made ofother smooth, electrically insulating materials such as a silicondioxide-coated silicon wafer, sapphire or certain ceramic and polymermaterials. The micro-beads may be made of an inert material such asborosilicate glass. The beads may also be made of other materials suchas latex and silicon dioxide.

The electrodes may be photolithographically fabricated on the membrane14. The generator electrode 18 is made of thin gold film over-coatedwith silver and a silver-chloride (AgCl) film. Chloride ions (Cl⁻) arethus generated as the spectator species by the application of a negative(reducing) current pulse to the generator electrode. The active layer ofthe sensor electrode 20 is also silver/silver-chloride (Ag/AgCl).

The diffusion of the chloride ions is monitored by recording the E vs ttransients at the sensor electrode 20. The rise and the shape of thetransients, as noted, depend on the free volume of the solution in thechannel between the generator electrode 18 and the sensor electrode 20.The rate of diffusion of the spectator species in the inter-electrodegap is thus used to measure the population density of the beads 22 inthe channel 16.

To incorporate the DNA recognition element of the sensor, the ssDNAstrands 24 a (the probe ssDNA) complementary to the target DNA arechemically bound to the channel surface in the inter-electrode gap. Theyare used as molecular tethers immobilizing the individual beads in thechannel 16. Each bead is attached to the device surface by a smallnumber (as low as one) of the dsDNA tethers 24.

As represented by FIG. 1C, the DNA sensing process involves competitivehybridization between the surface bound ssDNA fragment 24 a in theinter-electrode gap and the target DNA 25. This involves breaking ordehybridization of the dsDNA tether and results in the release of atleast some of the micro-beads. This is possible since the originaltether attaching the beads consists of a shorter section of the dsDNAthan the one formed during the sensing process.

Following the initial recording of the E vs t transient, the top plate12 of the device is removed and a small volume (as small as 1 nL) of ananalyte or sample solution is placed on the surface of the channel 16and then equilibrated. This could involve a period of 5 to 60 minutesand a slightly increased temperature above room temperature. Presence ofthe target DNA in the analyte solution will result in its hybridizationwith the complementary surface bound DNA 24 a, as mentioned above. Thiswill result in a release of at least some of the beads from the surface14 a of the plate 14. The term release means a complete release of abead from a channel surface. In some cases, not all the tethers joininga bead to a channel surface will be broken. However, in other cases,this will occur. In either case, the free volume of the analyte solutionin the channel will be increased.

Consequently, a second measurement of the spectator species diffusion inan aqueous solution in the channel, carried out after rinsing thesurface of the device (channel) and replacing the top plate to recreatethe channel, will accurately report the difference, if any, in thepopulation of the beads in the channel. The rinsing step is designed toa) rinse away the micro-beads that are no longer tethered to the devicesurface, b) to remove the analyte sample and thus to terminate theincubation process and c) to establish the exact and desired baseconcentration of the spectator chloride ions in the channel. (SeeEquation 1). The second measurement will measure the new populationdensity of the micro-beads in the channel. If the latter is smaller thanthe initial than this indicates that the specific, target ssDNA waspresent in the analyte. By quantitatively comparing such measurementswith the set of control experiments done with known concentrations ofdsDNA (sensor calibration), one can also deduce the concentration of thespecific ssDNA in the sample.

The top plate 12 may be removably secured to the channel sidewalls. Assuch, the plate 12 may simply be lifted up to introduce a solution intothe channel. Alternatively, the plate 12 may be fixed to the channelsidewalls, and the plate would include inlet and outlet ports for theintroduction and removal of an analyte solution and of the releasedbeads. This would require that the top plate have two positions: higherand lower so that the beads, should they become released, could then bewashed away. Suitable microfluidic control mechanisms such as valves,pumps and routers could be used to introduce and remove a solution fromthe channel.

The inter-electrode gap may also include electrically-conductive pads(not shown) located between the generator electrode 18 and the sensorelectrode 20. These pads may be formed by evaporating gold on one orboth of the surfaces 12 and 14. The pads would not be in electricalcontact with either the generator or sensor electrode, but wouldfunction as independent electrodes. Thus, the pads should be no closerthan about 1 micrometer to either the sensor or generator electrode.These pads would be used if it were necessary to bind different types ofssDNA to different devices of a larger array of sensors. The sensor 10,as discussed below (see also FIG. 2A), would also include an additionalcounterelectrode (described below). A common reference electrode wouldbe involved in applying and controlling the potential of the pad with anauxiliary potentiostat (not shown).

The sensor 10, as discussed below, includes a reference electrode and acounterelectrode. A potentiostat (now shown) is used to provide adesired potential relative to the reference electrode. Thecounterelectrode closes the potentiostatic circuit.

As shown in FIGS. 2A, 2B and 2C, a sensor device 40 includes a parallelarray of electrodes. Specifically, the sensor device 40 includes acentral generator electrode 1 8a positioned between two sensorelectrodes 20 a and 20 b. The generator electrode is located at thejuncture of the channels 16 a and 16 b of the device. Theinter-electrode gap (g) between the generator electrode 18 a and sensorelectrode 20 a can be the same as that between the generator electrode18 a and the sensor electrode 20 b. As such, the generator electrode issymmetrically positioned between the two sensor electrodes.

The target substance introduced into the channel 16 a may be the same asthe target substance introduced into the channel 16 b. Alternatively,the target substance in the respective channels may be different.

The sensor device 40, in this latter configuration, is designed tosimultaneously detect different target substances such as differentssDNA fragments of specific sequences. The biological binding mechanismor tether in the respective channels would also be different. Therefore,the DNA tethers are selected to be complementary to the different targetDNAs.

The sensor device 40 further includes two counterelectrodes 42 and 44.These electrodes are two symmetric, externally shorted bars. They assurean even current distribution. A reference electrode 46 is also part ofthe sensor 40.

As discussed, a galvanostat is used to apply a constant current betweenthe generator 18 a and counterelectrodes 42, 44. In addition, in orderto apply potential to the inter-electrode pads, if present, an auxiliarypotentiostatic circuit would be used. It would consist of the pad,functioning as a working electrode, and an additional counterelectrode(not shown in FIG. 2A). The common reference electrode 46 would be usedto complete the potentiostatic circuit. The potential of each pad of thearray of sensors would be independently controlled. A pad, however, andits counter and reference electrodes need not be positioned in anyspecific way relative to each other. This is because that circuit wouldnot pass essentially any current. As mentioned above, one referenceelectrode could function for both circuits of the generator electrodeand the conductive pad.

The sensor and generator electrodes are photolithographically fabricatedusing gold on a glass substrate. The electrodes may be fabricated byelectroplating of silver followed by its partial oxidation to AgCl in aKCl electrolyte. The reference and counter electrodes can also bephotolithographically fabricated. The area 48 (the diagonally hatchedpattern) is coated with an insulator, such as a polymer, to determinethe length (l) of the generator and sensor electrodes. The dotted linerectangular area 50 marks the area enclosed by the upper plate 12. FIG.2B illustrates the beads 22 attached to the channel surface in theinter-electrode gap. FIG. 2C shows the channel without the beads.

The beads, as noted, may act as spacers between the upper and lowersurfaces of the sensor. Alternatively, spacer dots, for example, polymerspacers, may be lithographically deposited at the corners of the upperor cover plate 12. The thickness of the spacers may be approximatelyequal to, for instance, the diameter of a spherical bead. A mechanicalsqueezing device can be used to assemble the sensor device.

The sensor device 40 includes two channels, and the associated sensorand generator electrodes. It is also possible to construct a device thathas more than two channels along with the necessary sensor and generatorelectrodes for testing whether the analyte solution contains more thantwo specific ssDNA fragments.

An array of sensors includes an ensemble of individual sensors describedabove. Each sensor would include generator and sensor electrodes, eachhaving their own inter-electrode gap, either including or not includingelectrically-conductive pads. The purpose of such an array would be tocontain a large number of sensors in one hand-held device which whenexposed to one aliquot of a sample would simultaneously indicate thepresence of several target compounds of interest.

As shown in FIG. 3, the E vs t transients due to chloride ion (Cl⁻)diffusion are reproducible. Specifically, three sets of chloride ionElectrochemical Time-of-Flight (ETOF) experiments (points) and threefits of Equation 1 to the data using channel thickness (h) as theadjustable parameter (lines) are illustrated. The experiments were donewith a device like that of FIG. 2A wherein g=50 μm, the average heighth=3.09 μm, i=1.0 μA, [Cl⁻]_(init)=1.00 mM, and D_(Cl) ⁻=1.90×10⁻⁵ cm²/s.However, the channels did not contain any beads. (See FIG. 2C).

The sensor response (E vs t) can be obtained using the Nernst equationwith the experimentally obtained slope value. Using 50 μm gap (g)devices with 3.1 μm thickness (h), a good agreement between the theoryand experiment was obtained, as shown in FIG. 3, for Cl⁻ diffusion. Todocument reproducibility of the device assembly, three transientsobtained in three separate experiments are shown involving repetitiveopening and reassembling of the device 40. The average thickness of thechannel obtained by fitting h as the variable is 3.09±0.02 μm (±0.6%).The rate of Cl⁻ generation can be varied in the range 0.01-1 μmol/s(i_(gen)=1−100 mA/cm²). The lower limit is set by the maximum time of 1ms allowed for the double-layer charging of the generator electrode.

There exist numerous well developed procedures for binding ssDNA toglass and other solid surfaces. For example, one strategy relies onbiotin-streptavidin binding. In the first step, glass beads are exposedto a solution of bovine serum albumin (BSA), a protein known to stronglyand irreversibly bind to glass surfaces forming a monolayer coating.Inclusion of a controlled fraction of biotinylated BSA will allowcontrol over the coverage of streptavidin and consequently the ssDNA.Five′- biotinylated DNA strands of the type 5′-biotin-(dT)₁₀-TGT GCT AGTACA GAC-3′ can be used. The 10 thymine segment functions as a spacergroup. The fifteen bases “recognition” segment can be used to bind thebeads to the probe DNA on the device surface. The 25 bases probe DNA canbe attached to the gold coated inter-electrode surface, assuming goldpads are present in the inter-electrode gap, via thiol chemisorption.This strategy, as opposed to BSA adsorption, is chosen for two reasons.BSA adsorption might interfere with the micro-electrode functioning.Furthermore, the ability to independently control the potential of thegold pads in the inter-electrode gap will be essential in thepreparation of the micro-array devices where individual units will carrydifferent probe DNA strands. Chemisorption of thiol DNA derivatives suchas 5′HS—(CH₂)₆-AGA TCA GTG CGT CTG TAC TAG CAC A -3′ is impeded at highnegative potentials. Therefore, keeping just one unit of an array atopen circuit and the rest at about −1.1V will direct assembly of aparticular probe DNA onto a single, selected inter-electrode pad. Thelast 15 bases of the probe DNA are complimentary to the terminal 15bases of the biotinylated bead DNA. The complementary target ssDNA maybe between 15 and 25 bases long. All custom synthesized ssDNA samplesare commercially available.

The number of DNA tethers that can be formed between a single bead andthe device surface depends on the surface concentration of the ssDNAfragments, the length of a tether and the radius of the bead. The latterdetermines the curvature of the bead surface. Relying on a purelygeometric argument, the area (S) of a spherical cap of height t (equalto the average length of a DNA tether) is S=2πRt; if t is estimated tobe 10 nm, then S=6.3×10⁻¹⁰ cm². The maximum surface density of theattached ssDNA chains on glass beads is limited by the size of BSA(about 44 nm²/molecule; streptavidin is smaller, assuming binding of onessDNA per biotinylated BSA). This gives 2.3×10¹² molecules/cm². Anestimate of the maximum coverage of about 2×10¹² molecules/cm² can beused as the maximum surface density of the double-stranded tethers.This, combined with the contact surface area S, yields 1000 as anestimate of the maximum number of tethers.

The force required to break or “melt” a double-stranded DNA tether isknown from the art describing single DNA molecule “stretching” methodssuch as “magnetic beads”, “laser tweezers” approaches and atomic forcemicroscopy. These measurements offer insight into the elastic andinelastic regimes encountered in such single molecule stretching and“unzipping”. In the latter case, when a 5′ and 3′ terminals of a doublehelix at one end of a DNA molecule are pulled apart, the forces areabout 10-15 pN (pico-Newtons) and correspond to stepwise breaking ofindividual base pairs (thus unzipping). When the opposite ends of adsDNA are pooled, the so called overstretching transition (also referredto as dsDNA melting transition) is observed in the 65 to 200 pN range.Much larger forces are involved in breaking a covalent bond withindsDNA. For example, a single molecule of dsDNA was broken by a recedingmeniscus method at a force of 960±140 pN.

In the sensor of the present invention, forces on the order of 100 pNwill likely be required to break a single tether. Thus, a beadattachment by even a single dsDNA tether is likely a sufficiently strongattachment to resist rinsing, thereby preventing false positivereadings.

The DNA surface density on glass beads can be determined byepi-fluorescence microscopy. Following DNA binding via theBSA/streptaviding scheme, the beads can be treated with a fluorescentlylabaled ssDNA reagent such as OliGreen (Molecular Probes Inc.). The goalis to control the binding density in the range of about 2×10¹²−2×10⁹molecules/cm². This spans the range of the attachment densities from ahigh of 1000 tethers per bead to one tether per bead. In the lattercase, a single bead would carry only about 200 DNA strands (and thusfluorophores). Nevertheless, these measurements are well within thesensitivity of the fluorescence microscope equipped with a cooled CCDcamera. The surface density of the DNA strands on gold will be measuredusing a chronocoulometric procedure. It involves saturation binding ofruthenium hexamine to the DNA phosphate groups. Precise DNA surfacedensity is obtained from the intercepts of the Anson plots after thedouble-layer capacitance charge subtraction. Sensitivity of thesemeasurements extends to about 1×10¹² molecules/cm². Clearly, in order tocontrol the number of dsDNA tethers binding individual beads to the goldsurface, the DNA binding density on the beads will need to be controlledvia control of the mole fraction of the biotinylated BSA in the BSAsolutions.

Following incubation of the DNA derivatized device surface with asuspension of the DNA derivatized beads and rinsing, the bead surfacedensity can be determined by optical microscopy. Determination of thenumber of tethers attaching each bead is far more challenging. However,it does not exceed the sensitivity of fluorescence microscopy with athermoelectrically cooled CCD camera. One approach is to rely on dimericcyanine dyes (such as YOYO-1 characterized by and available fromMolecular Probes Inc.) specifically developed to stain dsDNA. These dyesare known to fluoresce only after intercalating into dsDNA. To avoidquenching, gold surfaces must not be used in these measurements.Instead, probe DNA can be bound to glass surfaces viabiotin-streptavidin protocol. This will not significantly alter thenumber of tethers per bead as long as that number is determined by thesmaller surface density of the ssDNA on the beads.

The response of the chloride ETOF devices loaded with the beads needs tobe calibrated. Since regardless of the bead attachment density thesystem is always above percolation threshold, the shape of E v ttransients can be quantitatively predicted knowing the number of thebeads in the inter-electrode gap and the dimensions of the diffusionchannel. Thus, the effective porosity or void-volume fraction (ρ, seeEquation 1) can be calculated and compared with the experimentalresults. Any discrepancy between the experiment and the calculationswould most likely be due to the difference between the nominal andactual bead size or due to bead size dispersion. In either case, theseexperiments will allow for the calculation of the average bead diameterto be used in the subsequent DNA sensing measurements.

The sensor should be highly sensitive and have a reasonable responsetime, for instance, on the order of about 10-30 min. As discussed,target DNA sensing involves recording two E v t transients, before andafter exposure of the sensor to a DNA sample. Following sensorcalibration, the top plate of the sensor can be removed and a smallvolume of a sample solution carrying the target DNA deposited onto thesensing surface. Following incubation and rinsing, the top plate isreassembled and the E v t transient recorded.

The sensitivity of the sensor depends on a number of variables,including the concentration of the target DNA and the length of thetarget DNA sequence. As discussed, the beads are attached via DNAtethers. Each tether may contain a 15-base double-strand. The probe-DNAmay be 25 bases long. The length of the complimentary target DNA may bevaried between 15 and 25 bases, to optimize sensitivity. In view of thesmall size of an individual inter-electrode gap area, the volume of thesample solution could be exceedingly small (for example, about 1 nL).The sample solution should also carry a known chloride ion concentration(about 1.0×10⁻⁴ M). The later is an input parameter in theinterpretation of the ETOF transients (see Equation 1). Temperature andincubation time are also naturally important parameters to control.

Detection of the competitive probe-target DNA hybridization requiresrelease of at least one bead. This in turn requires dissociation of alldsDNA tethers linking a bead to the surface of the device. Due to theconstrained space between a bead and the device surface, the sensitivityof the sensor most probably is inversely related to the number oftethers binding individual beads.

A variable pressure flow cell can be used to develop a controlled way ofrinsing the device surface to remove the unattached beads followingprobe-target DNA hybridization. It will allow control over the waterflow velocity and thus the hydrodynamic force acting on the beadsattached to the cell's surface. Briefly, such a cell may consist of twoparallel glass plates (about 2×10 cm²) separated by a distance w=20−50μm. In addition to this parallel plate narrow slit element, the flowsystem can include a pump, a buffer reservoir and a flow meter. Themaximum water (buffer solution) velocity (ν_(max)) that stillcorresponds to laminar flow conditions can be calculated from thecritical value of a Reynolds number, Re≦100 for the narrow slit cell.Re=wpρν_(max)/μ, where ρ and μ are water density and viscosity. V_(max)may approximately equal 100 cm/s. Such flow will require a hydrostaticpressure (ΔP) of 4 bar (calculated using ΔP=8 μLν_(max)/w², where L=10cm, the length of the cell). Flow velocity of this magnitude would exerta force F_(max) of 1 nN (nano-Newton) on a 2 μm bead attached to one ofthe plates of the cell. The estimate of the critical Reynolds number israther conservative. It is likely that 5 to 10 times larger hydrodynamicforces can be generated with this flow system. This simple apparatuswill allow microscopic observation of the release of the beads followingincubation with a target DNA sample solution, as well as an assessmentof the minimum force required to detach the beads. Note that theestimate of the force required to “melt” a single dsDNA tether of about65-200 pN is well within the force range of this flow cell apparatus.Therefore, a distinction should be able to be made between beadsattached by 1 to 10 tethers just by controlling buffer flow velocity inthe parallel plate cell and relating it to the hydrodynamic forcerequired for detachment. This capability will be important in theprocess of optimizing sensor sensitivity. The estimates of thehydrodynamic forces in the flow cell also work under the conditions oflow surface densities of the beads.

The smallest number of detached beads that can be detected depends onseveral design parameters. In view of Equation 1, it is clear that thesensor response is sensitive to the changes of the fractionalfree-volume (or permeability ρ) in the channel. Changes on the order of5-10% can be easily detected visually. To optimize sensitivity, thedevice size can be minimized and thus the total number of the beads inthe inter-electrode gap is minimized. To increase sensitivity, it isalso possible to use micro-beads larger than 2 μm in diameter.

A more precise and practical approach to obtain channel permeability(and thus bead density) relies on the measurements of transit times, τ,defined as time passed between the application of the current pulse andthe sensor potential reaching a value, for example, 50 mV or 70 mVnegative relative to its initial value. From calculated E v ttransients, it is known that τ depends linearly on ρ. Thus, theprecision of the measurement of the bead density will depend on a) thedefinition of τ, b) the noise level in the E vs t recordings and c) theprecision of the device assembly. The latter is less than 1%

The measurements of sensor potential are also of high precision (lownoise) and the fact that an absolute level measurements is not requiredis also an advantage. Altogether, it is expected that a 1% precision insignal transduction will be achieved. This corresponds to detection of asingle bead detachment.

While the above assessment of sensor sensitivity with respect to thebead population is straightforward, it is more difficult to predictsensor sensitivity with respect to the concentration of the target DNA.To achieve and to accurately measure very low detection limits, on theorder of low pico-molar (pM, or 10⁻¹² M ) or femto-molar (fM, or 10⁻¹⁵M), the surface concentration of the probe DNA in the inter-electrodegap should be decreased to limit sequestering of the target DNAunrelated to the competitive hybridization and bead release.

The sensitivity of the sensor will depend on a number of parametersrelating to the device design and to the equilibria involved in thecompetitive hybridization between the target and probe DNA strands.These are expected to depend on the difference in the number ofcomplimentary bases in the probe-target double-strand and in the ssDNAsegment linking the beads to the surface. In any given case, it isexpected that the sensor will exhibit an on/off type behavior(detachment of all the beads) when exposed to samples of complimentarytarget DNA with concentrations well above the detection limit. In theconcentration range approaching the detection limit, it is expected thata fraction of the beads will be detached. The detection limit could besub-pM. Most importantly, it may depend on and improve with thedecreasing number of dsDNA tethers liking individual micro-beads to thedevice surface. The key advantage of the ETOF signal transductionmechanism, in addition to the fact that it will require only simpleelectronic measurement circuit, is its intrinsic sensitivity. It couldultimately allow detection of single molecule events by transducingbreaking of a dsDNA tether into a detachment of a bead deduced bymeasurements of the time required to observe a specific change of asensor potential.

A number of implementations and techniques have been described. However,it will be understood that various modifications may be made to thedescribed components and techniques. For example, advantageous resultsstill could be achieved if steps of the disclosed techniques wereperformed in a different order, or if components in the disclosed deviceare combined in a different manner, or replaced or supplemented by othercomponents.

For example, other biological binding and cleaving events, such ascompetitive antigen-antibody recognition or enzymatic cleaving of aspecific peptide sequence, can be relied on as the process responsiblefor the selective recognition of a target substance, triggering therelease of a surface-bound bead.

For instance, as shown in FIG. 4, a sensor 60 uses a number ofantibody-antigen-antibody tethers to attach the beads 22 to a sensorsurface 14 a. The sensor is designed to respond to a particular antigen62. The tethers 64 consist of a sandwich of two antibodies 66 a and 66 b(represented by Y-like symbols), and two antigen molecules 62 bridgingthe antibodies. In this type of a sensor, two different monoclonalantibodies (one bound to a sensor surface 14 a and the other to asurface of a bead 22) could be used, each targeting a different epitopeof the antigen. Alternatively, a combination of a monoclonal andpolyclonal antibody, or just one polyclonal antibody attached to boththe bead and device surfaces can also be used. The scheme of competitiveantibody antigen interaction (proceeding from left to right in thefigure) involves an assumption of facile kinetics of theantigen-antibody dissociation. A reverse scheme in which the targetantigen recognition step involves micro-bead attachment to the devicesurface can also be implemented as described below.

An example of a target immunogen is free prostate-specific antigen(PSA). In view of the importance of sensitive and specific detection ofPSA in the early detection of prostate and breast cancer, the rationalof this selection is apparent. In this case, monoclonal PSA antibodies(mAnti-PSA) will be first attached to the device surface and saturatedwith PSA. The bead attachment will involve exposure of the devicesurface to a suspension of micro-beads carrying polyclonal PSAantibodies (pAnti-PSA) on their surface. As in the DNA case, the sensingprocess will involve competitive antibody-antigen interactions betweenPSA in the analyte solution with mAnti-PSA—PSA—pAnti-PSA tethers leadingto the release of the micro-beads.

Additionally, in yet another embodiment, a sensor includes anantigen-antibody complex comprising a modified antigen exhibiting alower binding constant than the native antigen. The sensor relies onbreaking or cleaving of this pre-established antigen-antibody complexvia competitive interactions with an analyte, the native antigen. Thiswill result in a free energy driving force of the exchange step. Thesensor uses the dissociation of the antigen-antibody complex both as away of achieving sensor selectivity and specificity, and as its signaltransduction mechanism.

As shown in FIG. 5, such a sensor 70 includes a number of microspheres22 bound to a sensor surface 14 a by a number of molecular tethers 72incorporating a modified antigen-antibody complex. The tethers 72include a modified antigen 74 and an antibody 76. The antigen 74 has alower binding constant than a native or target antigen 78. The antibodymolecules 76 are initially immobilized on the device surface 14 a whilethe modified antigen protein 74 is attached to the surface of themicro-spheres 22 by, for example, a molecular spacer or linker 74 a.Formation of the antibody-antigen complex 72 links the micro-spheres tothe device surface. This sets the initial, active state of the sensor.

The analysis step involves exposure of the sensor device to a volume,for example, of a sample solution. If the latter contains a targetedimmunogenic protein its interactions with the sensor's antigen-antibodycomplex during an incubation period results in the exchange of theantigens and the breaking of a tether 72.

Two conditions are required to make this scheme practical: the exchangeinteractions between the original antigen-antibody complex and theanalyte antigen must be energetically favorable as well as kineticallyfacile. The use of a modified antigen with an approximately 2 orders ofmagnitude lower binding constant than a target antigen will meet bothrequirements. Such modification will not only make the exchangeequilibria with the analyte antigen thermodynamically favorable, but itwill also improve the exchange kinetics. Knowing that K=k_(on)/k_(off)(K is the equilibrium constant or binding constant of theantibody-antigen complex, k_(on) is the binding rate constant, andk_(off) is the dissociation rate constant), surface modification of theantigen will result in an increase of the dissociation rate constant.Thus, following a simple rinse, breakage of all tethers holding amicro-bead can be easily detected with an optical microscope.Alternatively, the release of micro-spheres can be used to generate anelectrical signal. Another approach relies on paramagnetic micro-spheresor beads; their detachment in a magnetic field generated by a stack ofsmall permanent magnets circumvents the rinsing step. The general designof this approach is shown in FIG. 6.

As shown in FIG. 6, the sensor 70 is positioned on a stage 80 of aninverted microscope assembly 82. The microscope assembly 82 includes alight source 84 and a charged coupled device (CCD) camera 86. Light fromthe light source is directed to an optics portion 88 of the microscopeassembly via a transreflector 90. The camera 86 also views the optics 88by way of the transreflector. A stack or group of permanent magnets 92is located above the sensor 70. The position of the magnets is selectedto generate a desired magnetic force on the paramagnetic beads 22.

The tethers 72, in one embodiment, may include a modified PSA antigenmolecule 74 and a PSA antibody 76. The target antigen 78 is also PSA.The molecular linker or spacer 74 a may be attached to a bead 22 via astreptavidin protein 75. The antibody 76 is also attached to the sensorsurface 14 a via a streptavidin protein 77 and a protein 78, such asBSA, that is used to coat non-biological surfaces. BSA, as noted,strongly absorbs on glass, plastic or other surfaces.

One attachment scheme involves commercially-available, for example, fromBangslabs or Dynabeads paramagnetic, streptavidin-coated beads. Next PSAis biotinylated. This can be easily accomplished, for example, bytreating it with one of any number of NHS-biotin derivatives, includingmolecules with spacer groups between the hydroxysuccinimide and biotinylends of the molecule for easier binding, and the hydroxysulfo-groupsallow the reactants to all be water soluble. An additional factorencouraging the use of molecular spacers is that multiple bonds holdingmicro-spheres together increase the strength of the tether only when theantigen molecule is placed at the end of a spacer group. The amount ofPSA attached to the beads may be measured using fluorescently labeledPSA and epi-fluorescence microscopy. Standard streptavidin-biotincomplex formation is used. This is shown in FIG. 6 where thestreptavidin protein 75 is adsorbed at the upper bead surface. Biotin isa very small molecule in comparison. It is symbolized by a dot 79 in oneof the four docking grooves of the streptavidin protein 75 or 77.

To consider the sensitivity of a sensor, reference is again made to FIG.5. The active surface of an individual sensor may comprise an area ofapproximately 50×50 μm². Thus, an area of 1 mm² may contain an array of100 or more individual micro-sensors each designed to detect a specificprotein or a DNA sequence. The active area of each sensor may bepopulated by a random array of approximately 100 micro-beads (about 1 μmin radius). Each micro-bead may be attached to the device surface byapproximately 100 or fewer molecular tethers each involving a complexbetween a modified antigen and its specific antibody or a dsDNA.

The expected high sensitivity of the sensor derives from two factorsintrinsic to its design. First, the signal transduction (involvingeither optical microscopy or generation of electrical signal)incorporates an intrinsic amplification element in that a small numberof single molecule events (dissociation of the molecular tethers) resultin an easily detectable macroscopic event—detachment of a relativelylarge particle. The second factor concerns the antigen exchangereaction. The latter can be represented by:Ab−Ag′+Ag→Ab−Ag+Ag′  (2)where Ab represents antibody, and Ag′ and Ag represent the modified andnative or unmodified antigen protein, respectively. It is important tonote that the each exchange equilibrium involves an immobilized Ab−Ag′complex and an Ag protein both confined within a small reaction volumebetween the surfaces of the micro-bead and the sensor surface. Due tothis confinement effect, reaction or equation (2) can be considered tobe a unimolecular process. This then leads to: $\begin{matrix}{K_{e} = {{\frac{p}{1 - p}\quad{and}\quad{thus}\text{:}\quad p} = \frac{K_{e}}{1 + K_{e}}}} & (3)\end{matrix}$where K_(e) is the equilibrium constant of reaction (2), and p is theprobability of a successful exchange (breaking of a tether). Inmacroscopic terms, the latter is synonymous with the reacting fractionof the initial population of Ab−Ag′+Ag in the confined reaction volume.Naturally, K_(e) is simply the ratio of the binding constants of thenative (K) and modified (K′) antigen: K_(e)═K/K′. A surface modificationof an antigen protein can result in at least two orders of magnitudedecrease of its binding constant yielding p of 0.99. In thermodynamicterms, the “confinement effect” refers to the entropy of mixing which isnot a component of the free energy of reaction (2) when the latter takesplace in a reaction volume much smaller than K or K′. Next the questionis asked: how different is p if equilibrium or reaction (2) isconsidered in a classical “non-confined” system? Then, the relationshipbetween p and K_(e) is straightforward: $\begin{matrix}{K_{e} = \frac{p^{2}}{\left( {1 - p} \right)^{2}}} & (4)\end{matrix}$This yields p=0.91 for K_(e) of 100. Therefore, whether the confinementeffect is significant, it is concluded that essentially every encounterbetween an analyte antigen and a modified antigen-antibody complex willresult in its dissociation. In other words, when K_(e) is about 100 orgreater, reaction (2) becomes nearly irreversible. Naturally, thisthermodynamic argument must be broadened by considering the kinetics ofthe exchange and specifically the magnitude of the dissociation rateconstant, k_(off). The latter can be estimated assuming that the bindingrate constant, k_(on), is diffusion limited. Its magnitude is about10⁷-10⁸ M⁻¹s⁻¹. The latter is based on the fact that the rate constantof bimolecular collisions of the molecules diffusing with a diffusionconstant (D) of about 10⁻⁶ cm²/s is about 10⁹ M⁻¹s⁻¹, and that onlyabout 1-10% of those collisions are fruitful. As mentioned above, adecrease of K by two orders of magnitude puts its upper bound value atabout 10¹⁰ M⁻¹. Therefore, the dissociation rate constant can beexpected to be greater than 10⁻³-10⁻² s⁻¹. In the case of PSA, forexample, which exhibits K values of about 10⁹ M⁻¹, the k_(off)(following PSA modification) is expected to be 1-10 s⁻¹.¹¹ These valuesdo not forecast a need for excessively long equilibration times in orderto comply with the exchange kinetics. Thus overall, a crucial postulateis arrived at that within this set of constraints the sensor will act asan integrating device, as each antigen that reacts with and breaks atether is effectively trapped for the duration of the experiment.Therefore, sensor sensitivity is expected to be a time-dependentfunction since in the range of low analyte concentrations, the sensor'sresponse will become diffusion controlled. Heretofore, no existingimmunoassay method has exhibited this property.

To expand on this unique characteristic, the sensitivity of a sensordescribed above may be estimated. The detachment of 5 micro-beads as aconservative minimum detection limit is considered. Assuming that, onaverage, each one is attached by 100 tethers, accumulation of 500antigen molecules is necessary to generate a measurable signal.Diffusion of antigen to the active area of the sensor can be modeled ashemi-spherical diffusion to a disk with a radius of about 30 μm. Underthese conditions, the diffusive flux J(r) is constant and can beexpressed by: $\begin{matrix}{{J(r)} = \frac{4\quad{DC}^{*}}{\pi\quad r}} & (5)\end{matrix}$where D is diffusion constant of the antigen (D≅1×10⁻⁶ cm²/s) andC*[mol/cm³] is its initial bulk concentration. Clearly, accumulation ofa certain number of moles (N/N_(A)) of antigen diffusing, at constantrate, to the sensors active region is directly proportional to time andthe initial antigen concentration:N/N _(A) =J(r)πr ² t=4DC*rt   (6)where N_(A) is the Avogadro number, t is time, and r is the radius ofthe sensor containing the micro-beads responding to, for example, aparticular antigenic protein. In one embodiment, the sensor radius (r)may be about 30 microns, containing l00 or 200 micro-beads. Thusallowing, for example, for a 15 min incubation time, yields a detectionlimit of 0.08 pM. A lower detection limit could be realized by: 1)decreasing the number of micro-spheres in the individual sensor activeregion (This would decrease the sensor radius (r), and thus increase theantigen flux to the sensor. It would also decrease the threshold numberof detached micro-beads (assumed to be 5 in the embodiment above)); 2)decreasing the number of molecular tethers attaching each micro-sphere,or 3) increasing the incubation time. Thus, a sub-fM detection limit isrealistic following proper sensor optimization. Finally, it is worthpointing out that the micro-sensor detection limit is independent of theantigen-antibody binding constant. In the calculation above, it was onlyassumed that the antigen modification can result in a hundred folddecrease of its natural binding constant.

Several known chemical reactions may be used to modify specific aminoacid residues as a means of altering the topography of the epitope ofPSA involved in binding with a specific anti-PSA IgG. (PSA is prostatespecific antigen—our test case, anti-PSA is one of the PSA antibodyproteins.) One of several hydrophilic amino acid residues could berandomly targeted assuming that some of them are present in the epitoperegion of the protein. (Epitope region is a fragment of a proteinsurface which is contacted by a complimentary region of its antibody toform the antigen-antibody complex.) Excessive modification, however,would destabilize antigen-antibody complex and defeat the purpose ofusing the modified PSA to form stable molecular tethers attachingmicro-beads to the device surface.

The strength of the antibody-antigen interaction is a result of manyfactors, including ion pairing, hydrogen bonding, dipolar andhydrophobic interactions. See, J. Janin et al., “The structure ofprotein-protein recognition sites”, J. Biol. Chem., 1990, 265,16027-16030, which is incorporated herein by reference. The latterinclude steric considerations and solvent effects. Ideally, a smallmodification is induced such as altering the topography of the proteinsurface that would, for example, decrease the extent of hydrophobicinteractions or disrupt one hydrogen bond. Indeed, a loss of a singlehydrogen bond would result in a decrease of the binding energy by ca 20kJ/mol. This corresponds to changing the equilibrium constant by twoorders of magnitude, as desired. On the other hand, a more significantalteration such as, for example, elimination of a salt bridge, if suchexists, would likely excessively decrease the binding energy.Consequently, while little is known about the exact surface structure ofthe PSA epitopal region, for example, chemical modifications thateliminate charged groups are not at the top of the list of possiblemodifications to be explored. Bearing these general considerations inmind, the list of possible modifications includes the following:

Methylation of lysine. This can be carried out by adding mMconcentrations of formaldelhyde and a reducing agent such as sodiumcyanoborohydride to a protein solution in a4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (Hepes) buffersolution. See, N. Jentoff et al., “Labeling of proteins by reductivemethylation using sodium cyanoborohydride”, J. Biol. Chem., 1979, 254,4359-4365, and R. H. Rice et al., “Stabilization of bovine trypsin byreductive methylation”, Biochim. Biophys. Acta, 1977, 492, 316-321,which are incorporated herein by reference. Other reducing agents, suchas sodium borohydride, have also been used, but the cyanoborohydridereaction is less harsh, and is less likely to significantly alter theoverall protein structure, for example by reducing disulfide bonds. Thisreaction is specific to lysine and the amine terminus of the protein. Itdoes not eliminate the positive charge on the protein surface. Finally,should the addition of a methyl group only slightly reduce the bindingconstant, the formaldehyde could be replaced with ethanal or propanal toadd a larger ethyl or propyl group to the amine.

Methylation of histidine. This can be done using the procedure describedby M. J. P. Dekker et al., “Identification of a second active siteresidue in escherichia coli I-threonine dehydrogenase: methylation ofhistidine-90 with methyl p-nitrobenzenesulfonate, Arch. Biochem.Biophys, 1995, 316, 413-420, which is incorporated herein by reference.Briefly, a few mM of methyl p-nitrobenzenesulfonate (MNBS) is added to aprotein solution in a 100 mM potassium phosphate pH 7 buffer. Thisresults in the preferential methylation of histidine residues. Similarto the case of lysine, should a methyl group prove to insufficientlylower the binding constant, ethyl or propyl groups could be triedinstead.

Modification of cysteine. This can be accomplished by following theprocedure outlined by J. F. Schindler et al., “Conversion of cysteinylresidues to unnatural amino acid analogues”, J. Protein Chem., 1996, 15,737-742, which is incorporated herein by reference. In this scheme,alkane chains of any length can be added to cysteine chains simply byreacting the 1-bromo or 1-iodo alkane of the desired length with theprotein in a pH 9.5 buffer.

Iodination of tyrosine. In view of the size of iodine, its modificationmight prove excessive. The effect of this modification on the bindingconstant must be experimentally determined. Even if the size of iodineis excessive, tyrosine modification could prove effective if thetyrosine is located near the edge of the epitopal region. lodination canbe accomplished by adding five- to ten-fold excess of I₂ dissolved in KIand 25 mM sodium borate to a solution of protein. This results in theaddition of 1-2 iodines per tyrosine. See, R. L. Lundblad, “ChemicalReagents for Protein Modification”, 3^(rd) Edition., CRC Press, BocaRaton, Fla. 2005 pp 213-215, which is incorporated herein by reference.

Addition of various groups to arginine. For example, when protein isadded to a pH 7.4, 0.125 M Tris buffer containing 0.1 M KCl, 0.01 MMgCl₂ and a dilute suspension of dinitrophenol, the dinitrophenol addsto any exposed arginine residues. See, H. M. Levy et al, “Inactivationof myosin by 2,4-dinitrophenol and protection by adenosine triphosphateand other phosphate compounds”, J. Biol. Chem., 1963, 238, 3654-3659,which is incorporated herein by reference. Alternatively, a 3% solutionof phenylglyoxal in buffer can be added to the buffered protein solutionand the reaction allowed to proceed for 24 hours. This results in theaddition of phenylglyoxal to arginine groups. See, K. Takahashi, “Thereaction of phenylglyoxal with arginine residues in proteins”, J. Biol.Chem., 1968, 243, 6171-6179, which is incorporated herein by reference.

Reactions involving carboxyl groups. These schemes commonly involveWoodward's reagent K being added to the protein in an ˜pH 3 solutionalong with a nucleophile such as ethyl amine, and results in theformation of an amide. See, R. B. Woodward et al, “A new synthesis ofpeptides”, J. Am. Chem. Soc., 1961, 83, 1010-1012, and P. Bodlaender etal, “The use of isoxazolium salt for carboxyl group modification inproteins”, Biochemistry, 1969, 8, 4941, which are incorporated byreference herein. This scheme would eliminate a negative charge on theprotein surface and might be too harsh. However, if a carboxyl groupresides near the edge of the epitopal region, its modification may leadto a desired effect.

Regarding a DNA version of a sensor, it should be appreciated that thedifference in the binding constants between a modified antigen and anantibody relative to native antigen-antibody complex is, in the DNAsensor 10 (FIGS. 1A-1B), intrinsically reflected in the different lengthof the two ssDNA fragments: one, longer bound to the sensor surface (therecognition DNA) and the other, shorter bound to a bead surface. Thetarget DNA is expected to be longer than the ssDNA bound to the bead. Asa result it will quickly be able to replace and break the tether 24.

In the case of the enzymatic cleaving of a specific peptide sequence, inanother embodiment, a molecular tether would be forwarded attaching thebeads to the surface using that particular peptide chain. When a samplecontains the enzyme that cleaves that peptide the bead is released.

The general idea of competitive interactions between a target compoundand the same compound forming a tether that attaches a micro-beadrequires lability of the target compound—recognizing compoundinteractions. In other words, slow kinetics of ssDNA-ssDNAhybridization/dehybridization or antigen-antibodyassociation/dissociation could require excessively long incubationtimes. While the kinetics of DNA hybridization has been shown to besufficiently facile to be compatible with this sensor's principle ofoperation, the kinetics of antigen-antibody binding has not beeninvestigated in many cases. The available data suggest that the range ofthe dissociation rate constants (k_(off)) exhibited by the variousantigen-antibody complexes spans several orders of magnitude from about10³ to 10⁻⁵ s³¹ ¹. Clearly, those antigens which exhibit k_(off) valuesof their antigen-antibody complexes smaller than about 10⁻² s⁻¹ couldnot be easily detected with the technique described above. In thosecases, an alternative strategy exists. It involves the reverse of thecompetitive binding approach discussed above.

For example, in the embodiment of FIG. 4, the device surface would bepopulated with the antibodies specific for a particular target antigen,but would not contain any micro-beads. Micro-beads would also carry ontheir surface a small number density of the antibody molecules for thetarget antigen. Initially, a small volume of the analyte solution wouldbe deposited onto the channel surface and incubated. The antigen, ifpresent, would bind to the antibodies on the sensor surface. Next, aftera rinse, a second incubation would expose the device to a suspension ofmicro-beads carrying on their surface monoclonal antibodies specific fora different part, epitope of the antigen, or a polyclonal antibody (amixture of antibodies specific for the same antigen but recognizing andbinding to its various fragments). This will result in attachment of thebeads to the device surface. Following a rinse, as discussed above, theE vs t transient would report the population density of the beads whichin turn would allow a determination to be made as to the presence andconcentration of the target antigen protein in the sample. The reversescheme, of course, could also be used for DNA and peptide sensing.

In a reverse approach, in the embodiment of FIGS. 5 and 6, the presenceof specific antibody could be determined. To accomplish this, a targetantibody would be first chemically modified to get Ab′. It would then beused to attach micro-beads to the sensor surface. The same attachmentprotocol would be used except native (unmodified) PSA antigen, forexample, would now be used. Now, exposure of a sensor to a small volumeof a serum sample would result in bead release if a PSA antibody ispresent. This is because that native, unmodified antibody would easilyreplace the modified antibody in the attaching tether and thus break thelinkage.

A generally applicable signal transduction mechanism designed tofunction in a hand-held biological sensor and sensor arrays has beendescribed. Two applications of the transduction mechanism have beenpresented. One application is targeting immunogens (antigens), a largegroup of macromolecules (for example cancer-specific proteins, toxinsand numerous other pathogens) capable of inducing a humoral immuneresponse, in other words, inducing the generation of antibodies. Theantigens must be poly or at least bivalent (exhibiting more than oneepitope, an active region on the surface of an immunogenic macromoleculeinvolved in selective binding with the antibody). Since essentially alllarge macromolecules exhibit this property, this class of biologicalsensors is designed to target a broad range of antigens. The secondapplication deals with DNA sensing. Indeed, in both of these areas, few,if any, generally applicable signal transduction schemes exist thatcould be relied on in sensor design. Indirect techniques such asenzyme-linked immunosorbent assay (ELIZA) and assays relying on thepolymerase chain reaction (PCR) for signal amplification are notcompatible with sensor methodology as they involve secondary reagentsand time consuming processes.

A sensor of the present invention uses a signal transduction mechanismthat is adaptable to a range of biological or chemical sensors. It cantranslate a chemical event, even a singular molecule chemical event,into an electrical readout. The signal transduction mechanism alsoincorporates a substantial amplification factor in that a singlemolecule event, such as DNA hybridization, could result in the releaseof a microscopic object, for example, a microbead.

Additionally, other strategies exist for determining the number of beadsreleased from the surface of the sensor channel relative to those thatremain bound to the channel surface. These strategies can be groupedinto two categories: (1) Those that can count the beads by measuringchanges in the volume of a solution confined within a sensor channeloccupied by the beads. These systems or methods require that the channelencompassing the beads be of the same dimensions before and after thechemical step resulting in bead release. (2) Those that count releasedbeads using a direct method.

The above-discussed signal transduction mechanism illustrated in FIGS.1A-2C is an example of the first group, that is, a system or techniqueinvolving measurements of the solution volume change in the sensorchannel. Other such system and techniques include conductivitymeasurements, coulometric measurements and spectrometric measurements.

Conductivity measurements are often used in liquid chromatographydetectors. The beads would be confined to a sensor channel of a fixedheight or thickness (h) with two metal, for example, gold, electrodesconstituting the opposing plates of the channel. The conductance of thesolution between the plates is proportional to the concentration of aninert electrolyte, for example KCl, in the channel; the spacing betweenthe electrodes; and the bead population density (the lower the beaddensity the higher the conductance of the electrolyte solution).

A technique involving coulometric measurements of a redox spectatorspecies in a solution in the sensor channel could also be used. Such asystem would require that one of the opposing plates of the sensorchannel be coated with a metal, such as gold, and serve as a workingelectrode. If the solution in the channel contained redox activemolecules of a known concentration, the total charge collected duringtheir electrochemical reduction or oxidation can be interpreted in termsof the population density of the beads in the sensor channel. One of anumber of redox active species could be selected for this purpose. Forexample, ferric ions (Fe³⁺), oxygen (O₂) naturally present in waterequilibrated with air, ascorbic acid (vitamin C), or chloride ions (Cl⁻,chloride is not directly electro-active but its content could bequantified by electro-oxidation of silver) could be used.

Spectrometric measurements of the absorbance due to a spectator speciesin the channel could also be employed. Here, the solution in the channelof the sensor would contain species absorbing light in a particularrange of wavelengths of the visible region. Numerous inexpensive, watersoluble, organic dye molecules, for example, such as those used as foodcoloring compounds, could be used. Transparent glass or polymeric beadswould be required in this case. A small light beam from a light emittingdiode or a similar light source would be directed to pass through thechannel in between and parallel to the plates creating the channel.Measurements of the light intensity with a small solid-state opticalsensor could be then interpreted in terms of the bead population in thesensor channel.

The second group involves the direct measurement of the bead populationin the sensor channel. For example, the embodiment of FIG. 6 uses anoptical microscope to detect bead detachment. Other optical measurementand electrochemical bead counting techniques can be used.

One optical measurement of the bead number density in the sensor channelis similar to the spectrometric measurement technique described above,except the beads would be made light absorbing. This could be done in anumber of ways such as by immobilization of light absorbing dyemolecules on the bead surface or by incorporating a light absorbingcompound into the bead forming material. Detection of the light passingthrough the sensor channel would indicate the number of the beads in thepath of the light beam. This measurement would not require that thechannel be of some known height.

The beads in the sensor channel could also be electrochemically counted.This system would require that the beads contain a magnetic core, forexample, one made of ferric oxide, and that they beelectrochemicaliy-active. Several different strategies could be used toaccomplish the latter. For example, the surface of the beads could becoated with a thin film carrying electrochemically active molecules.Alternatively, the beads could be over-coated with a thin film of ametal, such as silver. The latter could be converted electrochemicallyto a coating of AgCl. The sensor would comprise of two plates positionedparallel to each other at a distance larger than a bead diameter. Asmall permanent magnet would be positioned above the top plate. Thebottom surface of the top plate would be coated with, for example, gold,and would function as a working electrode. Any released beads would beattracted to the top plate by the magnetic force and thus makeelectrical contact with the electrode. The latter would be held at aconstant potential at which redox molecules on the bead surface or thesilver coating (in presence of chloride ions) would be electrochemicallyoxidized or reduced, generating a certain fixed electrical charge perbead. Thus, the total charge collected would be directly proportional tothe number of the beads released from the sensor surface.

Similarly, metal coated beads, for example, gold-coated beads, that donot carry any redox active molecules could also be used. Here, when thebeads contact the top plate electrode, which is held a constantpotential, a charge would flow due to electrical charging of the surfaceof the metal coated beads. This charge would be proportional to the beadsurface area and would likewise indicate its release form the sensorchannel surface. The total charge due to this electrochemical“double-layer charging” of the individual beads would be againproportional to the total number of the released beads.

In summary, the sensor strategy outlined above involves two chemicallyand physically independent processes: chemical sensing involvingbreaking of the molecular tethers used to attach the beads, and chemicalor physical detection of the resulting (decreased) bead population. Thelatter can be accomplished in a number of different ways as shown above.

In view of this general two step operation mode (chemical/biologicalsensing, and signal transduction generating an electrical, easy tomeasure signal), it is apparent to recognize that this entire scheme, asalso discussed above, could also be implemented in reverse order.Specifically, chemical/biological sensing could be reconfigured toresult in bead attachment, starting with beads attached to the devicesurface. The signal transduction would then report the increase of thebead population attached to the device surface using one of the severaltechniques discussed above.

Although many of the components and processes are described above in thesingular for convenience, it will be appreciated by one of skill in theart that multiple components and repeated processes can also be used topractice the techniques of the present invention.

While the invention has been particularly shown and described withreference to specific embodiments, it will also be understood by thoseskilled in the art that changes in the form and details of the disclosedembodiments may be made without departing from the spirit or scope ofthe invention. For example, the embodiments described above may beimplemented using a variety of materials. Therefore, the scope of theinvention should be determined with reference to the appended claims.

1. A sensor to determine if a target substance is present in an analytecomprising: molecular tethers attached to a sensor surface such that atleast some of a plurality of beads, as a result of a biologicalinteraction between a target substance and the tethers, can at least oneof be attached to the sensor surface by the tethers or released from thesurface by cleaving the tethers; and a measurement system configured andarranged to determine the population of beads at the sensor surfaceafter the introduction of the analyte into the sensor.
 2. The sensor ofclaim 1 wherein a tether includes a hybrid double-stranded DNA molecule,an antibody, a modified antibody, a modified antigen or a peptide.
 3. Asensor to determine if a target substance is present in an analytecomprising: a plurality of beads located in a channel and attached to asurface of the channel by a molecular tether that can be selectivelycleaved in the presence of a target substance to release at least someof the beads from the surface of the channel; and a measurement systemconfigured and arranged to determine the population of beads in thechannel after the introduction of the analyte into the channel.
 4. Thesensor of claim 3 wherein the target substance is a target antigen and amolecular tether includes a modified antigen and an antibody with themodified antigen having a lower binding constant than the targetantigen.
 5. The sensor of claim 4 wherein the modified antigen has anapproximately 2 orders of magnitude lower binding constant than thetarget antigen.
 6. The sensor of claim 3 wherein the target substance isa target DNA and a molecular tether is a hybridized double-stranded DNAtether with one of the single-stranded DNA fragments of thedouble-stranded DNA tether being complementary to the target DNA.
 7. Thesensor of claim 6 wherein the hybridized double-stranded DNA tetherincludes first and second single-strand DNA fragments, the firstsingle-strand DNA fragment being attached to a surface of the channel,the sequence of the first single-strand DNA fragment being complementaryto the target DNA, and the second single-strand DNA fragment beingattached to a surface of a bead.
 8. The sensor of claim 3 wherein themeasurement system is a signal transduction system, a conductivitymeasurement system, a coulometric measurement system, a spectrometricmeasurement system, an optical measurement system, or an electrochemicalcounting system.
 9. A sensor to detect a target antigen comprising: aplurality of beads attached to a sensor surface by a modified antigenand an antibody wherein the modified antigen has a lower bindingconstant than the target antigen.
 10. A sensor to detect a targetantibody comprising: a plurality of beads attached to a sensor surfaceby a modified antibody and an antigen wherein the modified antibody hasa lower binding constant than the target antibody.
 11. A sensor todetect a target substance in an analyte comprising: a channel; aplurality of beads located in the channel and attached to a surface ofthe channel by a molecular tether that can be selectively cleaved in thepresence of a target substance to release at least some of the beadsfrom the surface of the channel; measurement means for determining thepopulation of beads in the channel after the introduction of the analyteinto the channel.
 12. A method of detecting a target substancecomprising: providing a number of molecular tethers at a channel surfaceof a sensor such that at least some of a plurality of beads, as a resultof a biological interaction between a target substance and the tethers,can at least one of be attached to the channel surface by the tethers orreleased from the channel surface by cleaving the tethers; introducingan analyte into the channel; and thereafter, determining the populationof beads at the channel surface to indicate a presence of the targetsubstrate in the analyte.
 13. The method of claim 12 wherein the tethersinclude a hybrid double-stranded DNA molecule, an antibody, a modifiedantibody, a modified antigen or a peptide.
 14. The method of claim 12wherein the target substance is a target DNA and a molecular tether is ahybridized double-stranded DNA tether with one of the single-strandedDNA fragments of the double-stranded DNA tether being complementary tothe target DNA.
 15. The method of claim 14 wherein the hybridizeddouble-stranded DNA tether includes first and second single-strand DNAfragments, the first single-strand DNA fragment being attached to asurface of the channel, the sequence of the first single-strand DNAfragment being complementary to the target DNA, and the secondsingle-strand DNA fragment being attached to a surface of a bead. 16.The method of claim 12 wherein the target substance is a target antigenand a molecular tether includes a modified antigen and an antibody withthe modified antigen having a lower binding constant than the targetantigen.
 17. The method of claim 12 wherein the target substance is atarget antibody and a molecular tether includes a modified antibody andan antigen with the modified antibody having a lower binding constantthan the target antibody.